The present invention relates to an implantable infusion pump. More particularly, the present invention relates to an implantable infusion pump which includes a pressure regulator apparatus for producing a constant drive pressure from a more variable driving force exerted on a drug solution contained in a variable volume drug chamber of the implantable infusion pump.
Infusion pump designs rarely appeared in the medical literature until the 1950s. Most of these early infusion pumps were extracorporeal. One such device included a reciprocating air pump driven by an electric motor. Yet another design considered comprised a metal housing for a glass syringe and a compression chamber fed by a tank of nitrogen gas. Yet another such infusion pump included a motorized syringe pump which included an electric motor connected to the worm drive that moved a syringe plunger by a gear box. The gears were interchangeable such that replacement of the gears permitted different delivery rates. Yet another infusion pump included a syringe plunger driven by a rider on a threaded shaft. Numerous other designs were considered for these extracorporeal infusion pumps. P. D. W. Soden in his thesis entitled "A Methodical Design Study of Miniature Profusion Devices For Chemotherapy of Cancer of the Head and Neck", studied possible designs for producing a miniature profusion device to be carried by ambulating patients receiving chemotherapeutic treatment for cancer of the head and neck. Quoting from his thesis, "Approximately two million alternative design solutions were synthesized and recorded in compact matrix form on a `morphological chart`". One of the numerous design concepts mentioned by Soden for possible use with an extracorporeal infusion pump was the use of a small tubular arrangement containing an elastic metal bellows possibly constructed from preloaded disks so as to form a relatively small diaphragm in the tubular arrangement for exerting a fairly constant force on the drug solution being infused. Due to the size of the diaphram, this design provided for very little, if any, compensation for changes in atmostpheric pressure.
One of the earliest implantable infusion pumps intended for use in laboratory animals comprised a micro-injector comprising a compressed spring held away from a rubber-capped glass tube by a metal alloy disk with a low melting point. Admisistration of the injection was accomplished by placing the animal near the coils of high-frequency induction heater. Activation of the coils melted the alloy disk and the spring ejected infusate into the desired site in the animal. A second implantable infusion pump for the continuous infusion of drugs utilized the osmotic pressure developed by a saturated aqueous solution of Congo red dye against water as its power source. The infusion pump comprised a partially collapsed rubber compartment filled with Congo red dye separated from a second water compartment by a semi-permeable cellophane member. Expansion of the rubber compartment as the water moved by osmosis into the Conge red solution ejected the drug from the infusion pump.
Implantable infusion pumps were clinically introduced in 1975. Implantable infusion pumps currently in clinical use or in animal trials anticipating clinical studies in the near future, include vapor pressure powered pumps, peristaltic pumps, and pulsatile solenoid pumps. The vapor pressure powered pump was developed at the University of Minnesota and is described hereafter. The peristaltic pump generally comprises a flexible tube placed in a U-shaped chamber in contact with rollers that press against the tube with sufficient force to occlude the tube's lumen. The rollers are rotated by a motor. As the rotor turns and the rollers compress the lumen of the tube, fluid is moved toward an exit. The rollers and housing are arranged so that a second roller begins to squeeze the tube before the first disengaged, preventing backflow of the infusate. Sandia Laboratories, Siemens AG, and Medtronic, Inc. have developed implantable pumps with peristaltic pumping mechanisms. A pulsatile solenoid pump includes a solenoid driven reciprocating chamber with two check valves to move infusate from the reservoir out through the delivery catheter. Infusate is stored in a flexible metal diaphragm reservoir. Such a pump has been developed by Fischell and colleagues at Johns Hopkins University Applied Physics Laboratory and by the Pacesetter Corporation.
Much effort has been expended in developing external infusion devices which provide a steady pressure on the drug solution so as to provide a steady flow of drug solution to the patient. For example, U.S. Pat. Nos. 2,815,152 and 3,023,750 as well as French Pat. No. 1,314,002 are examples of such devices.
Currently available implantable infusion pumps also have difficulty in maintaining constant pressure as the volume of the drug solution in their drug chambers changes. Typically, the output flow of drug solution is regulated by external means, an example of which is illustrated in U.S. Pat. No. 4,299,220, or if passive flow restrictions are used to control the drug solution output, flow variation must be tolerated. The two ambient conditions that commonly cause flow variation are temperature and atmospheric pressure. In the vapor-pressure powered infusion pump disclosed in U.S. Pat. No. 3,731,681, both of these conditions cause the pressure differential between the drug chamber and the internal body pressure to change thereby causing a corresponding change in drug solution flow rate from the infusion pump into an infusion site in the body. In addition, the spring action of the metal bellows typically used to separate the drug solution from the two-phase fluid adds a variable force to the otherwise volume independent force exerted by the vapor pressure, thereby causing a steady, although predictable decline in flow rates as the drug chamber empties.
In many applications it is necessary to change the flow rate of the drug solution frequently, more frequently than can be done by changing the concentration by an empty-refill cycle on a constant flow rate infusion pump. Examples of such applications are: (1) the delivery of insulin to a brittle diabetic with no residual insulin production, (2) the delivery of a chemotherapeutic agent that has a strong dependence on biological timing, or (3) the delivery of a hormone that is timed to the natural rhythm of the body.
Infusion pumps; for example, U.S. Pat. Nos. 4,373,527 and 4,146,029, have been developed which utilize electronic controls that respond to transmitted electromagnetic signals and thus can be programed by a non-invasive procedure. The electronics in these infusion pumps work relatively well due to the availability of very complex, low powered integrated circuits. However, such infusion pumps have complex flow control components that must respond to the electronic signals. An approach commonly used is to have the flow control device provide an impulse of drug solution flow for every impulse of electrical signal from the electronic control circuit. By having very small (microliter) individual impulses and repeating them within the normal clearance time of an infused drug solution in the blood stream (e.g. one to ten minutes), an approximation of steady flow is obtained. This method is very flexible in that both steady flow and variable flow up to bolus doses can be delivered by a single flow control mechanism. However, the high cycle rate of the flow control mechanism increases the wear rate of the components, increases power losses in start and stop events, and increases probability of failure of some component. If a particular component has a finite failure rate per cycle, the mean time to failure decreases as the rate of cycling goes up. When a repeatedly cycled valve is used to produce a constant flow rate for an extended time period (several hours) there is an unnecessary hazard involved that would not be present if the same fixed flow rate were achieved by other means. If the fixed flow rate were known, a simple capillary tube could deliver that rate with only one cycle of valve open and fixed dose, rather than a hundred or so open dose cycles which might be required in an electronic impulse controlled system.
Typical systems employed in such electronically controlled infusion pumps include: (a) cyclic filling and emptying of a small drug accumulator with upstream and downstream valves; (b) an active piston pump with passive valves; and (c) miniature roller (peristaltic) pumps. In all three of these mechanisms, the drug solution storage chamber is passive and is held at a fixed pressure usually a little above atmospheric pressure in order to suppress bubble formation from dissolved air. The low pressure serves to reduce the potential hazard of an infusator leak. Accumulator systems use a higher drug chamber pressure to get positive filling cycles.
The above described electronically controlled infusion pumps have an unnecessarily wide dynamic range and response time for many applications. Moreover, they are complex, expensive and subject to failure. On the other hand, the fixed flow rate unfusion pump has been shown to provide adequate therapy for a range of disease states with no flow control for a given cycle. An infusion pump is required which provides a degree of drug solution flow control which is better than currently available infusion pumps of the constant flow design but which is less complex than that of the presently available electronically controlled unfusion pumps.
The present invention solves these and many other problems associated with currently existing infusion pumps.